1. Field of Invention
This invention pertains to magnetic resonance imaging, and particularly to a method of operation and structure of a front end of a magnetic resonance imaging (MRI) system.
2. Prior Art and Other Considerations
Magnetic resonance imaging (MRI) involves the transmission of RF signals of predetermined frequency (e.g., approximately 15 MHz in some machines) via an RF coil structure into an object-to-be-imaged (e.g., a human patient). A short time later, radio frequency NMR responses are received by the same or similar RF coil structure. As will be appreciated by those in the art, it is possible to derive imaging information from such RF responses.
In general, the RF portions of an MRI system comprises an RF unit, an RF coil, and an RF "front end." The RF coil and the RF front end are housed in a shielded room (e.g. , a gantry room) . The object-to-be-imaged is positioned in or otherwise proximate the RF coil inside the shielded room, while the MRI system operators station themselves outside the shielded room.
The RF unit both generates the RF signals transmitted to the object-to-be-imaged and receives and processes the MRI signals returned from the object. For an MRI system, an RF coil typically comprises a conductor (e.g., copper) suitably geometrically shaped to couple with a selected "image volume" portion of the human body. In addition, the RF coil includes its own unique matching circuit.
The RF unit is interfaced to an RF coil by the RF front end. In general, an RF front end of an MRI system has one or more channels for transmission and reception.
Historically, only one coil could be attached or remain attached at any given moment to each channel of the front end. For example, in one prior art configuration, a sole transmit only (Tx only) coil was connected to a transmit channel and a sole receive only (Rx only) coil was connected to a reception channel. In another prior art configuration, a sole transmit/receive coil was attached to the transmit channel, and an Rx only coil attached to the reception channel. The prior art RF front ends did not permit connection of a quadrature transmit coil to the transmit channel, but did permit connection of a quadrature Tx only coil at the reception channel.
In the early days of MRI utilization, there were very few types of coils, such as a non-QD (non-quadrature) head coil, a non-QD body coil, and a few surface coils. Then, as until now, technicians had to enter the shielded room to attach a desired one of the coils to the RF front end. For example, if the scanning regimen for a patient required images using a head coil and then a body coil, a technician would have to enter the shielded room after completion of imaging with the head coil in order to detach the head coil from the RF front end and to attach the body coil.
As MRI technology improved, more types of coils have been developed, including more anatomy-specific coils, for improving diagnostics. Coil shapes have proliferated, at least in part due to the fact that a "filling factor" (i.e., a ratio of patient volume occupied inside a Rx coil over the total Rx coil volume) affects the signal-to-noise ratio (SNR) of an Rx coil.
As mentioned above, imaging information is derived from the NMR RF responses received via an RF coil structure. One significant limiting factor on the quality of such images is the attainable signal-to-noise ratio (SNR) of the RF responses which must be detected and processed. The RF coil impedance is transformed by the use of matching circuits to an optimum impedance required by a low noise preamplifier.
The RF coil structure (also known as the "RF coil assembly" and "applicator") is typically housed within a static magnet (e.g., cryogenic magnet) and gradient coil structure where its load impedance may vary considerably (depending, for example, on the size and composition of the object being imaged and which is therefore coupled to the RF coil). MRI signal processing circuits (e.g., in the RF unit), including RF transmit/receive circuits, are located remotely from the RF coil structure and are connected thereto by a transmission line. The signal-to-noise ratio is optimized when the RF coil/applicator is resonated so that the impedance looking into the coil/applicator (e.g., the "load impedance") is "matched" or made equal to the complex conjugate transmission line impedance that connects the coil/applicator to the MRI processing equipment.
The load impedance of the coil/applicator will vary considerably depending upon the mass and composition of the material being imaged. In the past, efforts were made to obtain satisfactory performance (favorable SNR) by making an inconvenient internal impedance adjustment to the RF coil (i.e., to an RF tuning/matching circuit located at the coil site) each time a new object was to be imaged.
Harrison U.S. Pat. No. 4,827,219, commonly assigned herewith and incorporated by reference herein, discloses a remote tuning network for matching MRI RF coil impedance. The remote tuning network includes two adjustable impedances for achieving a matched RF impedance condition between an RF coil assembly, an RF transmission line interconnecting the RF coil assembly with the remote tuning network, and RF processing circuits. The remote tuning network may achieve matched RF impedance over, for example, a 3:1 VSWR (voltage standing wave ratio).
Advantageously, remote tuning units (RTUs) such as those disclosed in U.S. Pat. No. 4,827,219 are low cost. However, having a complex transfer function, the remote tuning units generally do not quickly tune high power RF coils. The RTUs are tuned with conventional computer-controlled tuning algorithms which may require thirty seconds or longer to tune a coil. These conventional computer-controlled tuning algorithms use a "hunt and peck" approach (similar to conventional manual RF antenna tuning/loading adjustments) to minimize the magnitude of a signal reflected from the coil (e.g., the "reflected signal").
Thus, the cumbersome prior art practice of repeatedly entering a shielded room to manually attach and detach coils and repositioning patients, as well as the prior art "hunt and peck" tuning approach, cost precious time. Actual imaging time constituted only a small portion of the time that a coil was loaded with a patient. The non-imaging time expended in tuning operations and manual connection/disconnection of coils hindered greatly the efficient use of expensive MRI equipment, and tended to exacerbate concerns of anxious patients.
The present invention provides an RF front end for a MRI system which readily accommodates, at any given moment, attachment of a plurality of RF coils including RF coils having differing types of matching circuits. The included novel front end facilitates gathering of diagnostic and other information. In part, this is facilitated by quickly and efficiently tuning an RF coil for an MRI system. Efficient and accurate calibration of RF coil tuning for an MRI system is also achieved.